Magnetic resonance imaging (MRI) is by now a well known and well documented phenomena. It is known that nuclei in a body naturally assume random orientations with respect to each other. Using MRI techniques, certain of these nuclei can be manipulated to cause measurable magnetic responses from within the body being imaged.
In MRI, the object to be imaged is placed into a generally homogeneous magnetic field H.sub.o (FIG. 1) created by an external static magnet source, shown in FIG. 5. The H.sub.o field causes the above-described nuclei (or some percentage of them) in the object to align themselves in the direction of H.sub.o. Because of their magnetic moment, these nuclei spin at a particular frequency called the Larmor frequency. In FIG. 1, the nucleus is modeled as an arrow (indicating its magnetic moment) spinning in alignment with the z-direction
If no other perturbation is introduced to the system, the nuclei continue to spin in alignment with each other due to the presence of the H.sub.o magnetic field. But, in MRI, the nucleus is perturbed by an RF pulse created by RF transmission coils (FIG. 5). This RF transmission causes the nucleus to tip (or nutate) at a certain angle, usually 90.degree., relative to the direction of the homogeneous magnetic field. Because the nucleus is originally spinning in the direction of the homogeneous magnetic field before application of the nutating RF pulse, the RF pulse causes it to nutate to a plane orthogonal to the direction of the magnetic field (the x-y plane) where it will precess, as shown in FIG. 2.
Once the RF pulse is removed, the precessing nucleus gradually relaxes back into alignment with the homogeneous magnetic field H.sub.o, as shown in FIG. 3. Since each nucleus can be conceptualized as a small magnet, the vector component of magnetism from the nuclei in the direction of the homogeneous magnetic field is zero when the nucleus is fully nutated 90.degree. (FIG. 2). As the nucleus relaxes back into alignment with the homogeneous magnetic field (FIG. 3), the magnetization vector component in the direction of the homogeneous magnetic field gradually increases to a maximum when the nucleus returns to the fully aligned position, with a time constant T.sub.1. The component of the magnetic vector in the x-y plane decreases with a time constant T.sub.2. As the nucleus relaxes from the 90.degree. position to the aligned position, it transmits an RF signal at the Larmor frequency as a result of these changing magnetic vector components.
The RF signal emitted by the relaxing nuclei can be received by an RF receiver coil, processed by control and image processing circuitry, and displayed as a processed image signal on a video display tube, as shown in FIG. 5.
Of course, a body to be imaged contains many such nuclei. To obtain MR image signals, one usually isolates smaller volumes or areas of the body for selective analysis. In two-dimensional Fourier Transform MRI, nuclei in the body are spatially encoded using gradient magnetic fields in the three orthogonal directions of the image volume. Traditionally, the x, y, and z axes defining the gradient magnetic field directions are referred to as the read, phase and slice directions, respectively. The slice gradient (G.sub.z) imposes a gradient magnetic field along the length of the body being imaged in order to select a slice of the body for data acquisition. When data acquisition is completed for one slice, another slice will be selected until the full volume desired to be imaged has been slice selected and imaged.
Once a slice is selected, the area of the slice can be viewed as a two-dimensional image defined by the so-called phase-direction (y) and the read-direction (x), as shown in FIG. 4. In the phase direction, another magnetic gradient coil (this one having a variable amplitude (G.sub.y)) is used to spatially encode the nuclei along the y-axis by the phase of the spins. As shown in FIG. 4, the spins of the nuclei along the y-axis have identical frequencies, but differ in phase, or amount of rotation, relative to one another. This is accomplished by temporarily altering the frequency of the spins in the y direction after the 90.degree. RF pulse nutates them. If the frequencies of spin are temporarily altered unevenly in one spatial direction by a variable amplitude magnetic gradient field, then when the gradient field is removed, the nuclei will return to the same Larmor frequency but with different spatial phase encoding. By then viewing the signal strength of nuclei at a particular phase, one can determine the density of nuclei at the spatial location corresponding to that particular phase.
A similar spatial encoding occurs by the so-called read direction magnetic gradient field (G.sub.x). This gradient field changes the frequency of rotation of the nuclei across a line in the x-axis direction just as the NMR image signal is read out by the RF receiving coil. The result, as shown in FIG. 4 for example, is a spatial difference in frequency (speed) of the spins in the x-direction just as the NMR signal is read. Thus in FIG. 4, the nuclei in the left-most portion of the x-axis are shown spinning slowly, while the nuclei in the right-most portion of the x-axis are shown spinning more rapidly. By taking the NMR signal with the frequency encoding, the resultant RF signal, when Fourier transformed, will yield a curve of amplitude versus frequency. And since the amplitude corresponds to nuclear density and the frequency corresponds to particular locations in the x-direction, the curve also represents nuclear density versus x-direction location, i.e., an image of nuclei in x-direction space.
As is known in the art, the spatial phase encoding in the y-direction corresponds to lines of resolution in the final image. That is, a greater number of phase encodings along the y-direction can produce a greater number of lines of resolution in the final image. Typically, MRI systems have 64, 128, or 256 lines of resolution, and thus a corresponding number of phase encodings along the y-axis. For each encoding that is to result in a line of resolution, separate NMR signal data must be collected. As is also known in the art, one method of collecting NMR data is to flip the spinning nuclei 180.degree. to cause an echo (spin echo) of the NMR signal. After a single 90.degree. nutation pulse, several NMR signals can be collected by imposing several consecutive 180.degree. pulses to cause several consecutive spin echoes. By changing the phase encoding between each consecutive 180.degree. pulse, one can obtain data for several lines of resolution using only a single 90.degree. nutation pulse. This process of using multiple flips following the original nutation pulse in order to obtain multiple spin echoes is referred to as fast spin echo sequencing.
Unfortunately, the number of 180.degree. nutation pulses that result in useful data is limited by T.sub.2. As described above with respect to FIG. 3, following the 90.degree. nutation pulse, the nuclei begin to relax back into alignment with the H.sub.o field. While relaxing (T.sub.2), the nuclei can be flipped several times (changing the phase encoding each time) in order to obtain several spin echo signals. But, once the nuclei relax for several T.sub.2 's, the signal to noise ratio of the resultant NMR signal is too small for the signal to be discriminated and the process must be started again with a new 90.degree. nutation pulse. The present invention is not limited to any particular number of spin echoes obtained from each 90.degree. nutation pulse, but for purposes of illustration only, seven echoes are used in some of the following examples.
FIG. 6 illustrates a typical NMR 2DFT (two-dimensional Fourier Transform) sequence. FIG. 6 illustrates the general timing of the RF transmission pulses (RF) from the RF transmitter coil shown in FIG. 5. It also shows the timing of the gradient coil pulses (G.sub.x,G.sub.y,G.sub.z) in the read, phase, and slice directions, respectively. Finally, FIG. 6 illustrates the resultant NMR signal received by the RF receiver coil shown in FIG. 5.
The RF transmission signal begins with a 90.degree. nutation pulse occurring at the same time the slice gradient (G.sub.z) is selecting the slice of nuclei to be imaged. As a result, the nuclei in the slice are nutated 90.degree. (into the position shown in FIG. 2), leaving the nuclei in neighboring slices in alignment with the H.sub.o field. The nuclei in the selected slice immediately begin relaxing back into alignment with the H.sub.o field, as described with respect to FIG. 3. Thereafter, the first 180.degree. RF pulse is applied followed by phase encoding pulse imposed by the phase encoding gradient field G.sub.y at the first line of resolution (k=+1). As a result of the above steps, a slice of nuclei has been selected and within that slice a line of resolution has been selected. After an identifiable time lapse, the first spin echo (NMR) will occur, during which, the read gradient G.sub.x is imposed for frequency encoding.
After the signal reading, the selected nuclei are re-phased by an equal or opposite phase gradient (in this case k=-1) following the readout pulse and before the next 180.degree. RF pulse.
The NMR signal is stored in the image processing circuitry (FIG. 5) until a full collection of NMR data is obtained for a full slice. Once the data for the slice is obtained, each NMR signal in the slice is plotted into a two-dimensional array of data points, which when Fourier transformed in two dimensions will yield an image of the nuclei density at each pixel location of the slice plane.
The remaining NMR signals in the slice are obtained as shown in FIG. 6. After the first NMR signal is received and the nuclei are re-phased, a new phase encoding is imposed. In the example of FIG. 6, seven NMR signals are obtained from each 90.degree. nutation pulse. To cover 256 lines of resolution, approximately 36 scans (90.degree. pulses) must be employed (36 scans.times.7 NMR signals per scan=252 lines of resolution). When the k-space does not get filled by a whole number of scans, an additional scan can be used (i.e., a 37th scan) or the zero fill signal can be used to fill in the missing scan lines.
In the example of FIG. 6, just after the second 180.degree. nutation pulse (following the first 90.degree. nutation pulse), the phase gradient G.sub.y phase encodes at the position k=+19. The read gradient G.sub.x is then applied, and the second spin echo (NMR) is obtained. Afterward, re-phasing occurs by G.sub.y imposing an equal and opposite signal at K=-19.
After the third 180.degree. pulse, the phase gradient selects the k=+37 position and the third NMR spin echo is obtained. Thereafter de-phasing occurs at k=-37. This continues for seven spin echoes, each having an associated phase encoding.
In FIG. 6, the seven echoes are obtained after k=+1, +19, +37, +55, +73, +91 and +109 phase encoding signals, respectively. When the seventh spin echo has been obtained, the signal strength of the NMR signal has decayed due to the relaxing of the nuclei back into alignment with the static magnetic field, H.sub.o. Accordingly, after the seventh NMR spin echo is obtained, the nuclei are permitted to relax for a predetermined period of time and then the next 90.degree. nutation pulse is applied for the same slice selection.
After the second 90.degree. nutation pulse, collection of NMR data continues for the slice selection. For example, in FIG. 6, phase encoding picks up at the k=+2, k=+20, k=+38, k=+56, k=+74, k=+92, and k=+110 positions. Again, after each echo signal is obtained an equal and opposite phase correction signal (k=-2, -20, -38, -56, -74, -92, and -110, respectively), is applied to zero sum the phase angle. In this example, the second scan increments each k-space phase encoding signal by one over the k-space phase encoding of the first scan. The incrementation over the 36 scans will change after the eighteenth scan (following phase encoding of k=+18, +36, +54, +72, +90, +108 and +126), in order to avoid overlap of the k-space. Thus, the nineteenth scan may decrement from the k-space of the first scan in order to cover the negative k-space. In other words, the nineteenth scan NMR signals may be obtained following phase encoding of k=-0, -18, -36, -54, -72, -90, and -108 (with equal and opposite phase correction signals of k=+0, +18, +36, +54, +72, +90, and +108 following the respective NMR signals). Of course, +1 or -1 incrementation is not required; other k-space scan sequences are equally plausible, provided the cumulative scans cover the full k-space (k=-128 to k=128, in FIG. 6) for each slice.
After the seven NMR readings following the phase encodings of the second scan, the NMR signal has again decayed below an acceptable level and a new 90.degree. nutation pulse is applied. Although not shown in FIG. 6, one can appreciate that this process will continue for slightly more than 36 scans to obtain 256 scan lines of data, by incrementing the phase encoding k-space for each scan line. A 37th scan line can be used to obtain the few additional scan lines of data needed to complete the 256 lines of resolution in this example or the additional scan lines can be obtained from the zero k-space data, as described above.
FIG. 7 shows a short-hand way of illustrating the scanning sequence shown in FIG. 6. The seven 180.degree. nutation pulses per 90.degree. nutation pulse yield seven spin echoes, which are identified in FIG. 7 (like FIG. 6) by a circled number. The 256 lines of resolution in k-space phase encoding are on the horizontal axis, from k=-128 on the left to k=+127 on the right TE refers to the periodicity of the 180.degree. nutations following the 90.degree. nutation (echo spacing). That is, the seven 180.degree. nutation pulses occur, as shown in FIG. 6 at a particular period, which corresponds to 1TE through 7TE in FIG. 7.
As shown in FIG. 7 (and described above with respect to FIG. 6), the sequence of phase encoding is such that, from the 36 scans, the NMR signals following the set of first 180.degree. pulses (immediately following the thirty-six 90.degree. pulses) will correspond to approximately the k=-17 through k=+18 phase encoding positions. The NMR signals from the set of second 180.degree. pulses following the thirty six 90.degree. pulses will fill approximately the k=-35 through k=-18 and k=+19 through k=+36 positions. Similarly, the NMR signals from the set of third 180.degree. pulses following the thirty six 90.degree. pulses will fill approximately the k=-53 through k=-36 and k=+37 through k=+54 positions, and so on for all seven sets of 180.degree. pulses for all thirty six scans.
FIG. 7 also illustrates how the seven echoes following the thirty six scans will not completely fill the k-space. The NMR signals from the seventh set of 180.degree. pulses following the thirty six 90.degree. pulses will fill the far edges of the k-space only to positions k=-125 and k=+126. Of course, in the FIG. 6 embodiment, the full k-space extends to k=-128 and k=+127. The negative and positive k-space edges thus are left with, respectively, three and one unfilled resolution lines. These can be filled by either taking another scan, or the zero fill signal can be used to fill the missing scan lines.
In FIG. 7, the first echoes (the set immediately following the 90.degree. pulse) will be taken earliest in the T2 decay period and thus will dominate apparent contrast. In FIG. 7, these first echoes occupy the k=-17 through k=+18 position, i.e., the center of k-space, and correspond to low spatial frequency data. Similarly, in FIGS. 5, 6 and 7, the latest echoes (the seventh echoes) have the weakest signal strength due to T2 decay and thus contribute to the high spatial frequency data. As a result, the signal in FIG. 7 behaves as if a low pass (spatial frequency) filter had been applied.
The signals need not be obtained in the fashion of FIGS. 6 and 7. FIG. 8 illustrates the data acquisition sequence in which the last echoes contribute to the center of k-space and the first echoes contribute to the ends of k-space. The result is a signal that behaves as if a high pass (spatial frequency) filter had been applied.
FIGS. 9 and 10 illustrate still further alternatives in which a balanced acquisition of early and late echoes (i.e., high and low signal strength signals) contribute to high encoding. In FIG. 9, the high signal encoding is in the negative half of k-space and in FIG. 10, the high signal encoding is in the positive half of k-space.
It has been previously discovered that the acquisition time needed to obtain data for a full k-space can be reduced by slightly less than one-half if one employs Half Fourier imaging. The Half Fourier concept is described in Avram, et al, Conjugation and Hybrid MR Imaging, Radiology 189:891 (1993) and in Feinberg, U.S. Pat. No. 4,728,893, Increased Signal-to-Noise in MRI Using Synthesized Conjugate Symmetric Data (1988), both of which are incorporated by reference into the present disclosure.
An illustration of how Half Fourier imaging works is shown in FIGS. 11 and 12. In FIG. 11, for example, the full k-space data is acquired using only four echoes from each 90.degree. nutation pulse. The four echoes are shown as circled numbers and contribute to phase encoding in the negative half of k-space. The zero k-space location and the half of signal number 4 lying in the positive k-space location are also obtained during the scans.
In FIG. 11, the highly encoded high signals contribute to the (negative) k-space positions and the late echo number 4 contribute to the central k-space positions. Because of the complex symmetry that exists between Fourier transform pairs in the negative and positive k-space, the missing data in the positive k-space can be obtained by conjugate symmetry from the three signals lying fully in the negative k-space. This means that one needs to gather only roughly one-half of the data used to construct the NMR image. This can save approximately one-half of the data acquisition time as shown in FIG. 11 where only four TE periods are needed to complete the NMR image compared to seven such periods in FIG. 8. Alternatively the extra available acquisition time can be used to produce more slices, or to generate more echoes to improve image resolution.
In FIG. 11, the missing data, which is obtained by conjugate symmetry, is shown as dashed lines in the positive k-space while the actually acquired data is shown as complete lines with circled numbers. FIG. 11 corresponds to FIG. 8 since it acts as a high pass filter, but improves on FIG. 8 since it decreases the data acquisition time.
FIG. 12 is identical to FIG. 11 except that it behaves as a low pass filter, as did the sequence depicted in FIG. 7. Again, the sequence in FIG. 12 reduces acquisition time through complex conjugation, as compared to the sequence of FIG. 7.
What is common to all of the Fast Spin Echo techniques discussed above with respect to FIGS. 6-12 is that different components of k-space are encoded by data modulated by different amounts of T2. In other words, the T2 response is different for high and low spatial frequency features in the image. The T2 effect is inherent to Fast Spin Echo techniques and thus the spatial encoding in k-space must be considered inherent as well, even though corrections can be made (for example, Chen, U.S. Pat. No. 5,517,122, T2 Restoration and Noise Suppression of Hybrid MR Images using Wiener and Linear Prediction Techniques, (1996)).